Conventional Optical Coherence Tomography (OCT) is an imaging technique that has found clinical use in ophthalmology, and further is being developed for a variety of endoscopic optical biopsy instruments. Conventional OCT provides three-dimensional images of living tissue with micron-scale spatial resolution in both the transverse and axial dimensions. The transverse resolution of OCT is obtained by conventional optical scanning, while the axial resolution derives from interferometry. In particular, in conventional OCT, a broadband light source is split into signal and reference beams, which are recombined and detected (e.g., by a photodetector) after the signal beam has interacted with (i.e., been reflected from) the tissue under examination. The broadband light source used in OCT is a classical-state light beam. Accordingly, for purposes of the present disclosure, a conventional OCT imaging technique employing a classical-state light source is referred to as “classical optical coherence tomography” (C-OCT).
FIG. 1 generally illustrates an exemplary configuration for a classical optical coherence tomography (C-OCT) implementation. In FIG. 1, the scanning process used to obtain transverse resolution of an imaged sample has been omitted for simplicity, as the salient concepts of interest in the present disclosure primarily are germane to axial resolution. In C-OCT, a broadband light source 50 produces classical-state light, which is divided into two beams, commonly referred to as a “signal” beam 52 and a “reference” beam 54. These signal and reference beams are in a joint classical state with a phase-insensitive correlation. Light sources 50 commonly employed for C-OCT generally have center wavelengths in the range of 590 nanometers to 1550 nanometers, bandwidths of approximately 40 to 90 nanometers, and power levels from approximately 1 mW to 100 mW. Perhaps the most common light source employed in C-OCT is a superluminescent diode (SLD). High resolution C-OCT systems use Ti:sapphire lasers with bandwidths as high as 200 nanometers at a center wavelength of approximately 800 nanometers and a power level of approximately 20 mW.
As illustrated in FIG. 1, the signal beam 52 is directed to a sample 56 to be imaged (e.g., a tissue sample) having an impulse response h(t), and the reference beam 54 is directed to a variable delay (having a delay denoted as T). The variable delay often is implemented with a moving mirror. In FIG. 1, for simplicity, the sample is shown as being irradiated by the signal beam and transmitting a beam 53 after interaction with the sample (the sample is being shown imaged in transmission), but it should be appreciated that sample imaging in C-OCT generally is done in reflection (i.e., light reflected from an irradiated sample is detected to provide the imaging information). After the signal beam has interacted with the sample and the reference beam is passed through a variable delay, the resulting beams 53 and 55 are combined in a 50/50 beam splitter 60 for detection via a measurement of second-order interference in a Michelson interferometer arrangement 62, which is followed by differential amplification in an amplifier 66 with gain G, and post-detection image processing by processor 68.
Recent work with non-classical (quantum) light has led to an OCT variant known as “quantum optical coherence tomography” (Q-OCT). FIG. 2 generally illustrates an exemplary configuration for such an implementation. In FIG. 2, as in FIG. 1, the scanning process used to obtain transverse resolution of an imaged sample has been omitted for simplicity, as the salient concepts of interest in the present disclosure primarily are germane to axial resolution. Similarly, in FIG. 2, the sample 56 again is shown for simplicity as being irradiated by a signal beam 72 and transmitting a beam 73 after interaction with the sample (the sample is being shown imaged in transmission), but it should be appreciated that sample imaging in Q-OCT generally is done in reflection.
The Q-OCT implementation illustrated in FIG. 2 requires a “twin beam” non-classical light source 70, which at present is provided by the process of spontaneous parametric downconversion (SPDC). In particular, a parametric downconverter is employed to generate a signal beam 72 and a reference beam 74 having a joint-quantum state and a phase-sensitive correlation. In the low-flux limit, this non-classical Gaussian state becomes a stream of individually detectable biphotons, which is the required light source output for Q-OCT. Initial experimental demonstrations of Q-OCT employed a type-1 phase-matched lithium iodate downconverter in the light source 70, in which the generated light has a center wavelength of approximately 800 nanometers, a bandwidth of approximately 40 nanometers, and power levels between 40 picowatts and 700 picowatts. After the signal beam 72 has interacted with the sample 56 and the reference beam 74 is passed through a variable delay 58, the resulting beams are combined on a 50/50 beam splitter 60. However, the phase-sensitive signal-reference correlation cannot be measured with a Michelson interferometer; accordingly, as shown in FIG. 2, the combined beams are detected via fourth-order interference using a Hong-Ou-Mandel (HOM) interferometer arrangement 82 that is followed by post-detection image processing by the processor 68.
For the same optical source bandwidth, Q-OCT offers a two-fold improvement in axial resolution over what is obtained with C-OCT. Moreover, Q-OCT is immune to axial resolution loss caused by group velocity dispersion in propagation to and from the sample depth that is under examination (i.e., even order dispersion is cancelled). These Q-OCT advantages, however, are counterbalanced by significant disadvantages. For example, with respect to the non-classical light source employed in Q-OCT, at present the flux in the twin-beams produced by parametric downconversion typically is quite low in comparison with what is obtained from the broadband classical light sources used in C-OCT. Thus, the measurement time needed to collect a Q-OCT image is significantly longer than that for a C-OCT image. A second disadvantage of Q-OCT arises from the requirement for photon-coincidence counting detection in a Hong-Ou-Mandel interferometer, which is significantly more difficult to do than the standard Michelson interferometer measurement used in C-OCT. As a result, Q-OCT is currently a laboratory curiosity, whereas C-OCT already is employed for various clinical uses (e.g., ophthalmology, endoscopic optical biopsy instruments). Because of the low-flux nature of the twin-beams used in Q-OCT, this technique is impractical for long distance operation (e.g., laser radar). C-OCT, on the other hand, can use bright sources of light and hence may be applicable over path lengths of kilometers or longer.